Note: Descriptions are shown in the official language in which they were submitted.
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OPTICAL VASCULAR FUNCTION IMAGING SYSTEM AND METHOD
FOR DETECTION AND DIAGNOSIS OF CANCEROUS TUMORS
BACKGROUND OF THE INVENTION
1. Field of the Invention
This invention relates generally to medical imaging systems and methods. More
particularly, it relates to an innovative optical vascular functional imaging
technology with
significantly improved image quality, sensitivity and specificity,
particularly useful in early
detection and diagnosis of cancerous tumors such as breast cancer.
2. Description of the Related Art
Early detection is key to lower mortality rates associated with breast cancer.
There is a
continuing need for a better cancer screening system that can provide accurate
early
detection of breast cancer in a safe, noninvasive, relatively inexpensive
manner. To lower
the number of unnecessary biopsies, improved diagnosis tools are also highly
desirable.
Currently, the standard screening modality for breast cancer is X-ray
mammography.
Unfortunately, X-ray mammography is less effective at detecting cancer in
younger
women's breasts, which are denser than those of older women. Moreover,
although the risk
of carcinogenesis resulting from X-ray mammography is relatively low, concerns
about
rislcs of exposure over many years of screening are valid. For these reasons,
other imaging
techniques are being used and studied to augment X-ray mammography, including
ultrasound, MRI, Tc-99m sestamibi scintimammography, and PET. These imaging
techniques are known in their respective fields and therefore are not further
described
herein for the sake of brevity.
Optical imaging techniques have also been explored. Optical imaging has many
advantages, for instance, it is noninvasive, has no ionizing radiation, and
requires no
painful compression, etc. Optical mammography was closely studied in the 1970
and -
1980s and proved to be inferior to X-ray mammography. The primary problem with
optical
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mammography is its spatial resolution. Optical mammography has a spatial
resolution of
0.5 to 1 cm, which means that blurring reduces contrast in smaller tumors.
U.S. Patent Application Publication No. 20050010114 by Porath, published on
01/13/2005,
entitled "OPTICAL MAMMOGRAPHY" attempts to address this problem by selectively
imaging planes of the breast utilizing non-ionizing radiation. Porath's non-
ionizing
radiation imaging system uses a special contact window located between
radiation detectors
and tissue being imaged and a camera focused on a depth of a slice to be
imaged.
Others have suggested administering, by injection or topological application,
patients with
contrast agents to reduce scattering. For example, U.S. Patent Application
Publication No.
20030157021 by I~laveness et al., published on 08/21/2003, entitled "LIGHT
IMAGING
CONTRAST AGENTS" proposes that contrast enhancement rnay be achieved in light
imaging methods by introducing particulate materials as scattering contrast
agents.
BRIEF SUMMARY OF THE INVENTION
During the process of angiogenesis, tumors develop abnormal vasculature, and
as a result,
cancerous tissue is often hypoxic, a condition that can be observed with
hemoglobin
oxygenation measurements. The present invention utilizes the endogenous
contrast afforded
by the spectroscopic properties of hemoglobin together with exogenous
vasoactive agents
to improve detection of cancerous tumors with differential/dynamic optical
imaging
techniques.
We have discovered that inhalation of oxygen (O2) and carbon dioxide (COZ) can
lead to
significant contrast for in vivo optical imaging. Using 02 and C02 as
vasoactive agents to
stimulate vascular changes has the additional advantage of being relatively
safe,
noninvasive, and requiring no injection or lengthy times between
administration and
imaging.
Using differential imaging with inspiratory contrast, our experimental results
show that the
additional contrast facilitates superior imaging quality than that of static
(conventional)
optical imaging. The increase in contrast between tumor (cancerous) and normal
(noncancerous) tissue is dramatic. We have observed up to a factor of two
variation in
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signal change. Taking advantage of this exogenous enhancement of the
endogenous
contrast due to oxy- and deoxyhemoglobin, the present invention provides clear
contrasting
images that would be particularly useful in early detection and diagnosis of
cancerous
tumors, potentially including breast cancer in women who are 40 or younger.
According to the invention, an imaging system acquires images through the
breast. Images
taken before and during inhalation of 02 or C02 are subtracted. An enhanced
optical
vascular functional (physiological) imaging system monitors abnormal
vasculature through
optical measurements on oxy- and deoxy-hemoglobin during inhalation of varying
levels of
02 and C02. Where applicable, enhanced data analysis procedures are utilized
to facilitate
the image analysis on the large amount of data acquired. In an embodiment, a
single optical
imaging system monitors both static and dynamic contrast mechanisms, thus
providing the
best possible sensitivity and specificity.
Compared with what is achievable with the physical image information provided
by x-rays,
the present invention provides more specific functional image information
particular useful
for early detection and diagnosis of breast cancer. By detecting tumors
generally missed on
x-ray mammography (false negative results), the present invention can reduce
the economic
and human cost associated with later detection of disease. By reducing the
number of false
positive diagnoses, it could also reduce the worry and economic cost of
unnecessary
biopsies.
Furthermore, because of the low cost of optical instrumentation, the present
invention could
be used in combination with x-ray mammography, which should provide greater
sensitivity
and specificity than x-rays alone. With the transition to digital x-ray
mammography, the
present invention can even share the same camera with an x-ray imaging system,
providing
excellent registration of two different modalities.
Other objects and advantages of the present invention will become apparent to
one skilled
in the art upon reading and understanding the preferred embodiments described
below with
reference to the following drawings.
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BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1(a) is a schematic diagram of an immersion imaging system.
FIG. 1(b) is a schematic diagram of an immersion imaging system adapted for
animals.
FIG. 2(a) is a static image of a mouse taken at 840 nm at 134 s after
administration of
carbogen.
FIG. 2(b) is the image from FIG. 2(a) with background subtracted.
FIG. 3 shows the temporal evolution of regions of the difference images at 780
mn.
FIG. 4 shows the temporal evolution of regions of the difference images at 840
mn.
FIG. 5 shows the temporal variation of relative changes in total hemoglobin
(top),
oxyhemoglobin (middle), and deoxyhemoglobin (bottom) during carbogen
inhalation. The tumor region is shown by the dashed line; the region on the
mouse torso away from tumor is shown by the solid line.
FIG.6 shows the temporal variation of relative changes in total 02 content
(oxyhemoglobin change, minus deoxyhemoglobin change) during carbogen
inhalation. The tumor region is shown by the dashed line; the region on the
mouse torso away from tumor is shown by the solid line.
FIG. 7 shows relative concentrations of oxyhemoglobin (a) and deoxyhemoglobin
(b)
concentrations at 140 s (100 s after carbogen administration).
FIG. 8 shows normalized eigen value spectrum.
FIG. 9 shows first two eigen images from principal component analysis.
FIG. 10 shows the temporal variation of the eigen image scaling factor.
FIG. 11 illustrates imaging of a human subject with immersion of the breast.
FIG. 12 illustrates imaging of a human subject with immersion and mild
compression.
FIG. 13 illustrates a form of 3-D data for differential vasoactive imaging.
DETAILED DESCRIPTION OF THE INVENTION
A primary goal of the invention is to develop reliable and yet inexpensive
technology to
improve sensitivity and specificity (lower false-negative and false-positive
rates) for early
breast cancer detection and diagnosis. We have achieved this goal with
enhanced functional
(physiological) optical imaging using a new type of contrast based on the
unusual vascular
function of tumors (atypical oxygenation improvement, atypical vasoactivity,
and blood
pooling).
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Another goal is to improve imaging through dense breasts where X-ray
mammography is
less successful. We have been investigating this differential vasoactive
optical imaging
(DVOI) approach in animal model studies. That work has demonstrated strong
contrast
between cancerous and noncancerous tissue during differential imaging in
rodents in
association with inhalation of OZ/C02 gas mixtures.
The contrast achieved by DVOI results from the vasculature in tumors and can
arise from
atypical oxygenation improvement, atypical vasoactivity, and blood pooling, as
monitored
by varying the levels of inspired 02 and C02. These differential vascular
function
measurements can be used to augment the cancer-specific static contrast
derived from 1)
elevated hemoglobin concentrations from angiogenesis and 2) reduced local
hemoglobin
oxygenation from tumor hypoxia.
A single DVOI system can monitor both static and dynamic contrast mechanisms,
thus
providing the best possible sensitivity and specificity from an optical
imaging system. COZ
and 02 are attractive contrast-enhancing agents because they are benign, safe
at appropriate
concentrations and inhalation periods and require no injection or lengthy
times between
administration and imaging.
Using these inspiratory contrast agents, we observed strong contrast between
images taken
before and during inhalation. We found that optical techniques can detect and
locate
picomole variations in chromophore concentrations over optical thiclcnesses
comparable to
those of the human breast. In the following sections, we describe how the
specificity of the
differential contrast available with the DVOI approach is sufficiently
significant to allow
tumor detection with higher sensitivity, even at the poor spatial resolution
available using
optical imaging through the human breast.
Advantages of using DVOI for breast imaging include functional imaging (i.e.,
imaging
that provides information on tissue state and function), inexpensive
instrumentation, and no
ionizing radiation. DVOI could prove useful as a primary screening modality.
Alternatively, it would be very useful as a secondary imaging modality to X-
ray imaging
for diagnosing, staging, or monitoring treatment of breast cancer. Because of
its simplicity
and low cost, DVOI can be efficiently incorporated into an X-ray or ultrasound
imaging
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system to provide functional information to complement the physical imaging of
these
modalities. DVOI may prove more effective in imaging dense breasts and may
reduce or
avoid the unpleasant or even painful compression used for X-ray mammography.
Optical Breast Imaging
As discussed above, the primary problem with optical mammography is spatial
resolution.
Optical mammography has a spatial resolution of 0.5 to 1 cm, which means that
blurring
reduces contrast in smaller tumors. This limitation can be overcome by
providing
functional imaging information.
Functional Optical Imaging
Whereas X-ray imaging primarily provides structural information, optical
spectroscopy
imaging can provide information both on structure and tissue function. For
example, optical
measurements at different wavelengths can indicate total hemoglobin content
and
oxygenation-functional information that is significant for breast cancer
detection. Tumor
angiogenesis typically leads to elevated local hemoglobin concentrations. In
addition,
tumors are often hypoxic, which can be observed optically as a decrease in
hemoglobin
oxygenation. Because tumors that are more hypoxic tend to be resistant to
radiotherapy and
chemotherapy and are more likely to be metastatic or invasive, the degree of
tumor hypoxia
can be used to guide treatment.
Tumor morphology also provides a source of contrast through variations in the
optical
scattering coefficient. The inventive system augments functional optical
imaging with
differential measurements related to tumor vascular function, taking advantage
of the full
range of available optical contrast. The broadest use of available contrast is
the most
effective for improving sensitivity and specificity.
The atypical characteristics of vasculature produced through tumor
angiogenesis provide
the scientific basis for the differential vasoactive optical imaging approach
disclosed herein.
The following articles, incorporated herein by reference, disclose information
related to
tumor angiogenesis: J. M. Brown and A. J: Giaccia, "The unique physiology of
solid
tumors: Opportunities (and problems) for cancer therapy," Cancer Res. 58, 1408-
1416
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(1990; and P. Carmeliet and R. K. Jain, "Angiogenesis in cancer and other
diseases,"
Nature 407, 249-257 (2000).
Blood vessels in tumors often exhibit distended capillaries with leaky walls
and sluggish
flow. These properties provide at least three types of contrast for optical
imaging in
conjunction with varying levels of inspired OZ and COZ. These types of
contrasts are due to
atypical oxygenation improvement, atypical vasoactivity, and blood pooling.
Because both
OZ and COZ are vasoactive, atypical tumor vasoactivity arising from
administration of
changing levels of these gases should provide strong imaging contrast. Tumor
vessels are
often contorted and leaky; thus, blood pooling in these vessels will delay
response to
oxygenation changes, providing another good contrast mechanism. Blood pooling
itself can
contribute to the atypical oxygenation improvement in tumors. However, our
experiments
indicate that atypical oxygenation improvement persists beyond the transient
response
caused by blood pooling.
Using functional optical imaging, the DVOI system disclosed herein can
reliably measure
the unusual vasculature in tumors. For example, by comparing hemoglobin
content before
and after carbogen is administered, opposing vasodilation and vasoconstriction
responses
after 15% C02 and ~5% OZ (carbogen) inspiration are readily detectable.
Similarly, the
changing response in tumor oxygenation after increased OZ administration is
easily
measured by monitoring hemoglobin oxygenation levels before and after the 02
level is
increased. Changes associated with blood pooling are observable in delayed
oxygenation
changes in the tumor. The DVOI approach could also incorporate quantitative
measurements of oxy- and deoxy-hemoglobin to improve overall sensitivity and
specificity.
DVOI very possibly can provide functional discrimination between benign and
malignant
lesions. Benign lesions tend to have rounded vasculature while malignant
lesions tend to be
more angular. Because the vasculature is different, it is likely that the
vascular response to
O~, and COZ will also be different.
There are additional motivations for examining differential contrast such as
that associated
with tumor vascular function. First, because the breast is highly
heterogeneous, comprising
the lobes (glandular tissue), fat, connective tissue, ducts, and supporting
vasculature, using
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a broader palette of contrast mechanisms should provide more specificity for
optical
imaging and help compensate for that heterogeneity. Second, the more
successful
noninvasive optical measurements (e.g., pulse oximetry, functional brain
imaging) are
differential or dynamic. Finally, recent theoretical work has demonstrated
improved results
using dynamic or differential optical imaging techniques, both of which rely
on changes in
optical contrast over time. In the following examples, we combine dynamic and
functional
measurements to obtain the best possible results.
EXAMPLES
Differential Vasoactive Optical Imaging System Setup for Animals
To monitor contrast for a range of tumor sizes and stages of development, we
performed
DVOI on and took noninvasive measurements from mice and rats. To replicate
tissue
thicknesses similar to those of the human breast, we partially immerse the
anesthetized
animals in liquid tissue phantoms that simulate the optical properties of
human breast
tissue. Although this approach does not allow for the effects of tissue
heterogeneity in the
breast, it is the most practical method for studying contrast without actually
using human
subjects. The measurements are noninvasive and thus can be readily repeated on
animals as
our instrumentation and methods are refmed/optimized.
FIG. 1(a) shows a continuous wave (CW) immersion imaging system 100 for
performing
DVOI with immersion. The system 100 comprises a near-infrared (NIR) light
source and a
camera, both of which are connected to a computer capable of analyzing image
data in
substantially real time. An immersion container is positioned between the
light source and
the camera for holding the imaging subject. In some embodiments, the light
source is made
up of an array of bright light emitting diodes (LEDs) and the camera is a
digital camera
with high sensitivity and high SNR.
As one skilled in the art will appreciate, the system can be readily
implemented in various
ways. FIG. 1(b) shows an exemplary system 110 adapted for animal model
studies. In a
specific embodiment, these LEDs emit near infrared (NIR) radiation with peak
intensities at
either 780 or 840 nm (Epitex L780-OlAU and Epitex 840-O1KSB, respectively).
Switching
between LED arrays enables measurements at different wavelengths and the
determination
of hemoglobin content and hemoglobin oxygenation. We increased light
throughput onto
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the imaging sensor by 20% by installing a large-aperture lens with high NIR
transmission
(JML Optics).
The NIR light source is directed at the sample immersion box, which contains
the study
animal in a heated (37° C), matching medium composed of water, inlc,
and submicrometer
polymer spheres (Ropaque from Rohm and Haas Company). This immersion medium
approximates the scattering and absorptive properties of the mouse tissue. The
front of the
immersion box is imaged onto the camera. Images at each individual wavelength
are then
collected, digitized (~-bit resolution), and sent to the computer for
analysis.
' The DVOI system can be readily implemented with a variety of suitable
cameras, for
example, the Dragonfly CCD (charge-coupled device) camera (Point Grey
Research), the
Pulnix TM-9701 CCD camera coupled to a Stanford Photonics Gen III image
intensifier,
and the ImagingSource DMI~ 3002 IR. Preferably, the system employs the digital
Dragonfly CCD camera because it offers a significant improvement in signal-to-
noise ratio
(SNR) over other video cameras. Although the Dragonfly has a lower absolute
sensitivity in
the NIR region compared with the other video cameras, it has lower read noise
and is
capable of longer exposure times (>60 s), which is important for imaging
thicker tissue
samples. More expensive cameras are available that provide superior
sensitivity and
sensitive area such the Retiga Exi manufactured by QImaging.
The compensation provided by immersing the animal (or at least the region of
interest) in a
tissue phantom improves image quality by removing changes in contrast
associated with
changes in tissue thickness and geometry, allowing better use of the dynamic
range of the
camera and providing more uniform illumination. When the match is good, the
tissue
almost disappears, and the image shows variations due to internal structure
and contrast,
which is what we want for i~ vivo imaging. The immersion medium serves to: (1)
allow
study of an effective tissue as thiclc as is typical for the human breast, and
(2) enhance
measurements by eliminating the effects of boundaries. Although the tissue
phantom lacks
the heterogeneity of the human breast, there is considerable heterogeneity in
the animal
itself.
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Tissue phantoms are prepared using our established methods, which are
disclosed in M.
Gerken and G. W. Faris, "Frequency-domain immersion technique for accurate
optical
property. measurements of turbid media," Opt. Lett. 24, 1726-1728 (1999); and
X. Wu, L.
Stinger, and G. W. Faris, "Determination of tissue properties by immersion in
a matched
scattering fluid," Proc. SPIE 2979, 300-306 (1997), both of which are
incorporated herein
by reference.
After an initial tissue phantom is prepared, an animal with a target region to
be imaged is
immersed between the source and collection fibers, the changes in amplitude
and phase are
measured, and the phantom composition is adjusted according to the optical
properties
determined from the immersion measurement. This process is repeated until the
optical
properties of the immersion medium and the imaged tissue agree to within a few
percent.
The thiclrness of the tissue phantoms is varied by inserting Plexiglas sheets
into the box
containing the tissue phantom for the CW measurements.
Animal Models
Human breast cancer cells (MDA 231) and mouse embryonic fibrosarcomas were
grown in
Dulbecco's minimum essential medium (DMEM) with glutamine and 10% fetal bovine
serum. The cells were harvested when they were 80% confluent, using 0.25%
trypsin. Cells
were injected subcutaneously on the dorsum of the female athymic nude mice
(approximately 23 g, Harlan Laboratories). Both cell lines were used at a
concentration of
2-3 million cells in 100 ~,1 of DMEM for each animal. The tumor volumes were
measured
twice weekly.
Animal Imaging
Imaging experiments were conducted on animals with tumor volumes of 500-1000
mm3.
We used two-four animals for each experiment. After being anesthetized with 40
mg/lcg of
pentobarbital, the mice were secured to a 3-mm Plexiglas platform with black
vinyl tape.
Anesthesia was given in further doses of 20 mg/kg as needed to reduce stress
associated
with immersion and to keep the animal immobilized. Carbogen or air was
administered to
the immersed mouse via a nose cone at a flow rate of approximately 3 1/min.
The optical
path length of the immersion box was adjusted to match the thickness of the
mouse (~2-2.5
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cm). At this thickness, the exposure time of the camera allowed us to measure
both
wavelengths at approximately three frames per second.
Images of individual mice were recorded before, during, and after the
administration of
carbogen. FIG. 2(a) shows one of these static images taken 134 s following the
administration of the carbogen. The approximate outlines of both the mouse and
the tumor
have been placed on top of the image as a guide. The mouse's head is out of
the immersion
medium and is above the field of view. The hind legs and tail are seen at the
bottom of the
image. FIG. 2(b) shows this same image after the subtraction of a background,
which is
simply an image of the mouse before the carbogen was turned on. Although the
boundaries
of the mouse and tumor are obscured by the good match with the immersion
medium, it is
clear from the difference image in FIG. 2(b) that there are distinct regions
of contrast
between the tumor and the surrounding tissues of the mouse.
Temporal T~ariatioya itz Diffenesztial CofZtrast
The enhanced contrast between the tumor tissue and the mouse tissue due to the
inhalation
of the carbogen was monitored by averaging the changes in intensity over areas
within the
difference images. FIGS. 3 and 4 show these averaged data for differences in
the 780 nm
and 840 nm images, respectively. The squares represent changes in the tumor
tissue, the
circles indicate an adjacent region within the mouse that does not contain the
tumor, and the
line represents the average of a part of the image not containing the mouse.
The maximum change for both wavelengths is approximately X10 units, and it is
clear from
the figures that distinct differences occur for the dynamics of the tumor
tissue when
compared with the normal mouse tissue. Furthermore, the background, which is a
measure
of lower limits for detection, varies just X0.2 units.
FIGS. 3 and 4 indicate that several regions (e.g., near 55 s at 780 mn, and
near 135 s at 840
nm) show strong contrast between tumor and surrounding tissue. Additional
contrast is
found after the carbogen is stopped; for 840 nm, the relative intensity of
tumor and
surrounding tissue reverses.
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Although images at a single wavelength such as FIG. 2(b) can be useful for
cancer
detection, it is also of interest to determine the changes in oxyhemoglobin
and
deoxyhemoglobin. We have analyzed the same image data set used to produce
FIGS. 3 and
4 to calculate approximate path-integrated oxyhemoglobin and deoxyhemoglobin.
The
absorption at 780 nm and 840 nm can be described as:
,uQ = 2.3~H6CHb~+ eHboz [HbOz ~j (1)
where l is the wavelength of interest, [Hb] and [Hb02] are the concentrations
(moles/L) of
deoxygenated and oxygenated hemoglobin, respectively, and a is the molar
absorption
coefficient. Using Beer's Law, we can describe the change in the absorption
coefficient ma,
at time t after a baseline image has been talcen as:
~~p = ~p,r - ~p,baseline = 2.31oglo Ibaseline /l (2)
It
where I is the intensity of transmitted light and l is the pathlength in cm,
corrected
appropriately for the differential pathlength factor for the animal tissue. We
can obtain a
rough measure of the change in path-integrated oxyhemoglobin and
deoxyhemoglobin
concentrations by assuming that the differential pathlength factor is the same
at both
wavelengths. By manipulating equations (1) and (2), we see that:
O,upBO 2~3 * EHb ~HboZ ~[Hb'
840 840 840 3
4~'a l ~Hb ~Hbo2 d[Hbo2
Because of the finite bandwidth of the LEDs, we calculated the absorption
coefficient by
integrating the wavelength-dependent absorption coefficient with the
normalized spectra of
the LEDs for each wavelength respectively:
(4)
This led to the following equations for the concentrations of Hb, Hb02,
Hbtorar at time t:
hso Is4o
0[Hb~t) = 2.3 * 7.507 * 10-4 * loglo hso - 5.271 * 10-4 * loglo I84o lh
t
hso Is4o (6)
0[Hb02 ~t) = 2.3 * - 5.255 * 10-4 * logo h$o + 7.996 * 10-4 * logo I84o ll,
(7)
v[HbtOt~t) = v[Hb~t) + v[Hbo2 fit)
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We used these calculations to determine the approximate temporal variation of
the total
hemoglobin, oxyhemoglobin, and deoxyhemoglobin shown in FIG. 5. These values
were in
turn used to calculate the approximate change in 02 content (oxyhemoglobin
change, minus
deoxyhemoglobin change) shown in FIG. 6. Several observations arise from these
images:
The tumor vasculature shows more erratic behavior, as seen from the
oscillations at the
beginning of carbogen inhalation. The failure to return to baseline for the
total hemoglobin
concentration (FIG. 5), and the overshoot in OZ content at the end of the
carbogen
inhalation (FIG. 6). The magnitude in changes of oxyhemoglobin and
deoxyhemoglobin are
accentuated in the tumor (FIG. 5). The increase in 02 content of the tumor is
delayed
relative to the rest of the animal (FIG. 5 middle and FIG. 6), which may be
due to blood
pooling in the tumor.
The same processing used for FIGS. 5 and 6 can be used to produce images
representing
approximate path-integrated oxyhemoglobin and deoxyhemoglobin as shown in
FIGS. 7(a)
and 7(b), respectively. These differential vasoactive images show a dramatic
increase in
tumor contrast as compared with a raw or static image, see, e.g., FIG. 2(a).
Principal Compo~zent Analysis
The imaging experiments described above generated large sets of data.
Typically, images
with 105 pixels at two wavelengths are recorded every 2-10 seconds over the
cycling period
of carbogen administration (approximately 10 to 20 minutes). Based on these
experimental
results, we expect to see < 7% change in image intensity following carbogen
administration. Because extracting such small signal changes from large data
sets poses a
formidable challenge, researchers have developed techniques that generate
smaller sets of
orthogonal images to describe the generated data, see, e.g., incorporated
herein by
reference, L. Sirovich and E. Kaplan, "Analysis methods for optical imaging,"
in Methods
for 1~ Yivo Optical Imagihg of tlae Centf°al Nej°vous System, R.
Frostig, Ed. (CRC Press,
2001 ); and L. Sirovich and R. Everson, "Management and analysis of large
scientific
datasets," Intl. J. Supercomputer Applications 6, 50-68 (1992). In practice,
these methods
have been shown to accurately describe data sets of 10,000 images with only
100 eigen
images.
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In the most basic adaptation of these methods, known as principal component
analysis
(PCA), the set of recorded images is represented by:
f = f (t~ x) (s)
where x describes the spatial pixel grayscale values of the image, and t is
the time at which
the image data was collected. Researchers have shown that these images,
f(t,x), can be
decomposed into the set of orthogonal functions a,2(t) andjn(x) by:
f (t~ x) _ ~ ~" a,~ (t)~P" (x)
A series of T time images containing P pixels can be described by the matrix:
. f (1,1) . f (1,2) . . . f (1, P)
M = f (2~1) f (2~2) . . . f (2, p)
( 10)
f (T,1) . . . .. . f (T, P)
This matrix can then be decomposed into the different an(t) and jn(x)
components through
the general technique of singular value decomposition:
a" (1) cp" (1) ,u, 0
A" _ . , V" _ . , and U = . (11)
a" (T ) ~P" (p) 0
and
M=AUVt ( 12)
The columns of V contain the orthonormal spatial basis functions, the
orthonormal columns
of A describe the time-dependence of the spatial basis functions, and U
contains the
weighting factors for the two matrixes A and V.
As a first step in processing the data, we apply this simplified PCA method to
determine
changes in oxyhemoglobin and deoxyhemoglobin, scaled by some pathlength factor
l as
described above. The time-dependent images that describe D[Hb] and D[Hb 02]
were
ordered into a matrix as shown in equation (10), and the singular value
decomposition was
carried out to obtain the matrices A, U, and V . FIG. 8 presents a plot of the
normalized
scaling factors contained along the diagonal of U. Only the first three or
four eigen images
contribute significantly to the set of images that describe the_ hemoglobin
dynamics in our
study.
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FIG. 9 shows the first two eigen images corresponding to the first two columns
of matrix V.
The contrast between the tumor and the surrounding tissue is evident in the
second image.
The time-dependent weighting of the second eigen image in the D[Hb](t) and
D[Hb02](t)
sets of images can be determined from the matrix product of A~U, and is shown
in FIG. 10.
Differential Vasoactive Optical Inzagiug System Setup for Humans
DVOI is very effective for breast cancer detection, and is preferred for
screening young
women with lcnown propensity for developing breast cancer. Combined with
another
imaging modality such as x-ray imaging, the DVOI system can prove to be a
powerful tool
in combating the disease.
In the case of human subjects, different imaging methods may be used for
differential
vasoactive imaging of the breast. The imaging may be performed with or without
compression and with or without immersion. In some cases, optimal imaging
entails using
at least mild compression and immersion.
Mild compression is advantageous for two reasons: first, with compression the
total
imaging distance is less, leading to a higher SNR, and hence increasing the
likelihood of
detecting a smaller tumor. Second, X-ray mammography uses compression. The
combination of optical imaging with X-ray imaging provides a further
embodiment of the
invention-given the low-cost of X-ray imaging and the possibility that both
imaging
techniques could be performed simultaneously. In a still further embodiment,
both imaging
systems share the same detector in the case where digital mammography is used
via
semiconductor-based cameras. That combination would lead to an improvement in
sensitivity and specificity over either modality alone. This embodiment
requires
coregistration of images from the two modalities, which could be achieved most
practically
if compression is used. .
Preferably, immersion is used to achieve highest possible sensitivity of the
imaging. With
immersion, all portions of the breast are imaged, with nearly the same
illumination reaching
the detector and providing more optimal use of the dynamic range of the
camera. That is,
the entire image may be acquired with a high level of illumination, and hence
high SNR.
For the non-immersed breast, variations in the transmitted light intensity
across the breast
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will be large. To avoid camera saturation in the thinnest regions, low light
levels will be
obtained in the thicker regions. Thus, the thicker regions will have a lower
SNR, and worse
imaging results. Researchers have used the phase measurement available with
frequency
domain measurements to perform correction for edge effects. Immersion achieves
a similar
goal.
Immersion can be achieved in at least two ways as shown in FIGS. 11 and 12. In
FIG. 11, a
human subject lies prone on a table similar to a stereotactic breast biopsy
table with the
breast immersed in a matching medium below. The light does not have to pass
through the
entire human torso. The optical measurements can be made with the light
passing through
the region of interest only. In this example, the light source illuminates
across the breast
only and not the entire torso. Preferably, the subject is provided with one or
more premixed
gas mixtures containing vasoactive substanceslagents, such as oxygen and
carbon dioxide,
by any method and apparatus that conveniently and comfortably deliver the gas
to be
inhaled. The system set up in both FIGS. 11 and 12 is similar to those shown
in FIG. 1(a)
and FIG. 1(b), although the system set up shown in FIG. 11 can also be used
without
immersion.
In FIG. 12, the breast is surrounded with a doughnut-shaped transparent bag
containing a
tissue phantom liquid. The bag would be filled to a slight overpressure to
press against the
breast in a manner similar to a blood pressure cuff, except that the
overpressure would be
much less. This method would achieve the same advantage of immersion but with
less
preparation and cleanup required. Preferably, the second immersion method is
employed
where a new bag with fresh immersion medium is used for each human subject.
The
immersion mediiun should be maintained at 37°C.
Where possible, optical imaging is preferably performed before any biopsy
procedure. This
avoids any influence the biopsy procedure might have on imaging measurement
and
interpretation. The imaging may be performed using only one or two inhalation
protocols
so that the total imaging takes only a few minutes.
As one skilled in the art will appreciate, the relative sensitivity and
specificity of a
diagnostic method depend on the criteria used. Relevant criteria include
percentage change
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in hemoglobin content and hemoglobin oxygenation, and the relative signs
(i.e., did each
increase or decrease). By varying the criteria used either sensitivity or
specificity can be
made high, but at the expense of the other dimension. To assist in the
analysis of the data,
we use a receiver operating characteristic (ROC) curve, which plots
sensitivity versus false
positive fraction; the free parameter is the criterion or threshold used for
diagnosis. The
area under the ROC curve gives a measure of the quality of the method; an area
near 1 is
desirable. The ROC curves are prepared for each contrast mechanism and for the
contrast
mechanisms in conjunction.
Gas Protocols
Because of the different respiratory rate, heart rate, size, and the fact that
the animals used
in animal models of the invention are anesthetized and humans would not be,
gas protocols
are different for humans and animals. Measurements are performed on animals
andlor
humans with varying inhalation gas composition and administration time to
establish
proper protocols for gas inhalation.
In the examples disclosed herein, gas mixtures of air, 02, C02, and 02+C02 are
produced
on demand using computer-controlled gas flow controllers. In some embodiments,
two
gases axe used: OZ and COz. In some embodiment, three gases are used to
produce these
mixtures: nitrogen, 02, and COZ. In some cases, mixtures of these gases may be
prepared at
fixed mixture ratios, and the gas inhalation protocol would involve switching
between
breathing of the premixed gases.
The gas flow controllers can rapidly alternate among gas compositions,
continuously
varying the levels of COZ and 02 in, for example, a nitrogen buffer, or create
carbogen.
Because CO2 and 02 have opposing effects on vasculature (vasodilation versus
vasoconstriction, respectively), using these two mechanisms in opposition or
in alternation
should produce useful results from the differential vasoactive imaging. For
example,
elevated C02 levels may be administered for a period of one minute, followed
rapidly by a
period of elevated 02. The same protocol could be repeated with a small
overlap between
the elevated C02 and OZ levels.
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Carbon dioxide is toxic when administered at high concentrations and carbon
dioxide
levels must be maintained at levels of 5% or less to avoid such toxicity. The
literature
indicates that concentrations as low as 2% achieve practical vascular activity
for
radiotherapy with good patient tolerance, see, e.g., incorporated herein by
reference, H.
Baddeley, P. M. Brodrick, N. J. Taylor, M. O. Abdelatti, L. C. Jordan, A. S.
Vasudevan, H.
Phillips, M. I. Saunders, and P. J. Hoslcin, "Gas exchange parameters in
radiotherapy
patients during breathing of 2%, 3.5% and 5% carbogen gas mixtures," Br. J.
Radiol. 73,
1100-1104 (2000). In the present invention, carbon dioxide levels are
preferably 0% to 5%.
With the computer-controlled flow controllers, we can sequentially administer
different gas
mixtures, which may or may not be premixed, to the same individual, taking
care that the
vasculature recovers sufficiently between the changes. We expect more
effective
discrimination between cancerous and noncancerous tissue, which ultimately may
be
utilized as a means of distinguishing among different tumor types.
Image Atzalysis Tools
The measurements acquired for differential vasoactive imaging comprise three-
dimensional
(3-D) datasets as illustrated in FIG. 13. The two spatial dimensions and one
temporal
dimension differ from other 3-D imaging modalities such as MRI or computed
tomography
(CT), which have three spatial dimensions. For those imaging modalities,
visualization
tools often create 2-D images as cross sections through the 3-D data set.
Regions of interest
can be probed by changing the orientation of the cross section. This is
similar to an
ultrasound technician changing the orientation of the ultrasound probe.
Visualization of data such as in FIG. 13 can be performed by taking cross
sections at
different orientations. FIGS. 2-4 are examples of a cross section and a line
section through
such a data set at constant time and position, respectively. However, both the
spatial pattern
(such as FIG. 2(b)) and the temporal pattern (such as FIG. 3) are necessary to
define
features in this data set. Simultaneously capturing both of these features
requires a different
sort of image analysis tool. One such tool is PCA. Applying PCA, the DVOI
approach can
be readily adapted to allow automated data processing of temporal image data
sets of
oxyhemoglobin, deoxyhemoglobin, and total hemoglobin, and change in 02
content.
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To improve the results, the DVOI approach may also adapt methods such as
spatial and
temporal averaging conditioned on the image features and the use of a py~ior~i
information
such as the temporal profile of the gas inhalation protocol. For breast
imaging, the number
of eigen images may be larger. The DVOI approach may therefore adapt methods
for
classifying the eigen images (e.g., by tumor type, other feature such as blood
vessels).
Ezzlzazzcizzg the DiffeYezztial ~asoactive Optical Izzzagifzg System
As one skilled in the art will appreciate, the DVOI system disclosed herein
can be
optimized or otherwise modified to improve its performance by, for example,
adding
another wavelength to enhance the imaging of water, increasing the
illumination power, and
increasing camera sensitivity. These modifications can enable imaging through
large tissue
phantoms with SNR (signal to noise ratio) limited only by shot noise, which is
a
fundamental limitation for any imaging process. High SNR can be very effective
for
differential imaging because image heterogeneity is removed during the image
subtraction
process. That is, subtraction of two images taken of the same field of view
yields an image
of zero intensity if nothing has changed.
1. Enhance imaging of water
Water concentrations are known to influence measurements of hemoglobin. Thus,
performing imaging at a wavelength dominated by water absorption should assist
in
quantifying oxyhemoglobin and deoxyhemoglobin measurements. Because of the
high
fraction of water in blood, images with dominant water absorption should also
help monitor
blood volume directly. Although the change in water content associated with
vasodilation
or vasoconstriction is relatively small, we have found that the differential
imaging is quite
sensitive to such changes. Thus, it is possible to monitor changes in blood
volume directly
using differential images at 970 nm, a wavelength dominated by water
absorption. A water-
based measurement of blood volume can also provide information on blood plasma
changes, which are somewhat different from the changes provided by hemoglobin
measurements. Measuring blood plasma and/or monitoring blood volume changes
with
water absorption are not critical to the success of our imaging approach, but
they
potentially could make the overall imaging approach more powerful.
2. Increase illumination
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The images shown with reference to the Working Examples section were obtained
using 21
LEDs at each wavelength. The power available from the LED array can be
increased by a
factor of 20 with more LEDs. Their brightness can also be increased by
operating them at
higher drive currents. Burn-in tests showed that the LEDs can be operated
significantly
above their typical operating currents for many weeks without incurring
problems. The
LEDs are turned on for only short periods during imaging at each wavelength,
thereby
increasing the practicality of higher current operation without LED damage.
3. Increase camera sensitivity
The camera sensitivity can be readily increased with a more sensitive camera
such as a
Retiga EXi camera produced by Q-Imaging. This CCD camera is approximately two
times
more sensitive in the NIR than the one used in the Examples above. In
addition, the
camera-sensitive area is four times larger. These two improvements will lead
to an overall
enhancement in camera sensitivity of roughly a factor of ~. In combination,
the increased
illumination and more sensitive camera should improve overall system
sensitivity by more
than 100 times.
Although the present invention and its advantages have been described in
detail, it should
be understood that the present invention is not limited to or defined by what
is shown or ,,~
described herein. Known methods, systems, or components may be discussed
without
giving details, so to avoid obscuring the principles of the invention. As it
will be
appreciated by one of ordinary skill in the art, various changes,
substitutions, and
alternations could be made or otherwise implemented without departing from the
principles
of the present invention. Accordingly, examples and drawings disclosed herein
are for
purposes of illustrating a preferred embodiments) of the present invention and
are not to be
construed as limiting the present invention. Rather, the scope of the present
invention
should be determined by the following claims and their legal equivalents.
